MRI scanners have long facilitated non-invasive, high-resolution imaging of internal structures of the human body based on the principle of nuclear magnetic resonance (NMR). Generally speaking, under that principle, atomic nuclei of tissue to be imaged absorb and reemit applied radio-frequency (RF) radiation based on the resonant radian frequency (labeled the Larmor frequency, f0) with which the quantum spin of the nuclei precesses in an external polarizing, static magnetic field (referred to as B0). B0 is typically generated via a main coil, often referred to as the “magnet,” and is further altered by way of one or more gradient coils that vary the magnitude of the static magnetic field B0, and thus the resonant frequency f0, in space and time to allow for selective excitation of tissue. Generally, the static magnetic field B0 is aligned along the longitudinal axis of the bore of the MRI scanner into which subjects to be scanned, such as human patients, are placed.
To facilitate the actual scanning, an RF excitation coil may be used to apply an RF excitation magnetic field B1 orthogonally to the static magnetic field B0 periodically, resulting in a realignment of the spins of the atomic nuclei. After each excitation, a relaxation of the realignment results in an echo RF signal being emitted by the nuclei, which is then captured by the scanner to generate the image of the tissue. Depending on the particular MRI scanner, the echo RF signal may be received by the excitation coil, or by a separate RF detector coil or other receiver structure.
In many high-field (e.g., B0=3 Tesla (T)) MRI scanners, a “birdcage” RF coil, typically formed by two circular metallic loops (“end rings”) at opposing ends of, and transverse to, the MRI scanner bore, and interconnected by an even number of longitudinal straight metallic segments (“legs”), is used to generate the excitation field B1. Typically, lumped capacitors are located along the rings between each pair of legs to tune the structure to render a near-homogeneous distribution for the excitation field B1. This structure is often driven by two excitation ports located along one ring in time-phase quadrature (90 degrees out of phase with respect to each other), resulting in the generation of a circularly polarized (CP), and more specifically a right-hand circularly polarized (RCP), excitation field B1+ in the near-field of the coil that helps maximize coupling between the excitation field and the tissue nuclei spins, resulting in a higher-resolution scan.
Generally, the higher the static magnetic field B0 that is employed, the more sensitive the MRI scanner is in detecting tissue differences, especially tissues located deep within a patient. However, since the Larmor frequency f0 is proportional to the magnitude of the static magnetic field B0, the use of ultra-high-field (e.g., B0=7 T) MRI scanners results in a significant shortening of the RF wavelengths employed. Consequently, the use of quasi-static, near-field RF coils, such as the one described above, tend to become limited to smaller, body-part MRI scanners, as opposed to whole-body MRI scanners, as the magnitude of the static field B0 increases due to the resulting reduction in size of the near field.
To compensate for the near-field reduction, MRI scanner technology has been developed that employs RF excitation using far-field antennas generating traveling waves (TW) within the MRI scanner bore. Examples of such RF excitation far-field antennas include, for example, axially-placed double-loop coils, dipole arrays, and axially-placed helical antennas located some distance from the patient. Such technology, however, has typically resulted in a number of issues, such as a high specific absorption rate (SAR) in tissues located near the antenna, resulting in unwanted heating of those tissues, as well as rapid attenuation of the excitation field and the resulting echo response as distance from the antenna increases.
With the above concepts in mind, as well as others not explicitly discussed herein, various embodiments of systems and methods employing a subject-loaded helical-antenna RF coil are disclosed herein.